Silicon has begun to receive increasing attention for use in biomedical applications. In particular crystalline silicon has been utilized as a textured surface to guide cell alignment, to encapsulate cells for implantation, and as an electroactive substrate to stimulate excitable cells. Several properties of silicon have led to its use in these diverse applications: (1) well-described silane chemistries for immobilization of adhesive ligands, (2) wet and dry micromachining capability to form 3-dimensional structures on biologically relevant length scales, and (3) semiconductor properties that allow incorporation of microelectronic elements. In comparison, porous silicon, a nanocrystalline material generated by etching of crystalline silicon in hydrofluoric acid, has been less extensively utilized for biomedical applications. Its open pore structure and large surface area, combined with unique properties such as photo and electroluminescence have provided a platform for sensors for non-biological species (e.g. solvents, gases, and explosives) as well as biological species (DNA, proteins). Indeed, the range of tunable pore sizes (5 to 1200 nm) in porous silicon spans a range of sizes important in biology; small DNA fragment is on the order of a few tens of nm, proteins are generally in the 100 nm range, and bacteria and cells can be a few microns in diameter.
Previously it has been shown that manipulations of cellular microenvironment by “micropatterning” on inorganic surfaces can alter the behavior of cells in culture (Chen et al, 1997; Bhatia et al, 1999). Methods to alter the support for cell growth to allow for distinct localized cell adhesion involved the manipulation of glass, gold or polymer supports such that cell adhesion molecules were differentially deposited on the support (e.g. U.S. Pat. Nos. 6,004,444; 6,103,479 and 6,133,030; all incorporated herein by reference). A common method involves the use of photoresist, a UV-sensitive polymer. Borisilicate substrates (e.g. coverslips) are coated with photoresist and exposed to light through a mask, creating a photoresist pattern. Patterned substrates are used to control subsequent immobilization of extracellular matrix components (ECM) (e.g. collagen I). The localization of a specific ECM component allows for the adhesion of cells to specific regions of the substrate (e.g. primary hepatocytes adhere to collagen, but not to glass). In some cases, co-cultures of two cell types are achieved by subsequent addition of a second cell type to attach to the periphery. Thus, micropatterned arrays have been used to generate defined co-cultures of hepatocytes and fibroblasts for the study of the maintenance of cell fate and function (Bhatia et al, 1999). Similar arrays have also been used for use in an apparatus for cell based screening. The method may be used for the establishment of any of a number of patterns, including non-uniform arrays (U.S. Pat. No. 6,103,479). However, such a method requires that the cell types of interest have different adhesive properties that are well known. Thus the system is limited to the use of cell types with well defined, and distinct, characteristics.
Some researchers have begun to explore the use of porous silicon as a biodegradable material for the slow release of drugs or essential trace elements to cells or as an in vivo diagnostic [10-12]. Promising findings by Canham et al. have shown hydroxyapatite nucleation on porous silicon in vitro, suggesting that porous silicon, in contrast to crystalline silicon, could be a bioactive surface (Canham, 1995; Canham et al, 1997). Nonetheless, porous silicon has not been extensively characterized as a material for implantation or the formation of hybrid (biological/non-biological) devices in vitro (Rosengren et al, 2000). Studies on the compatibility of this material with mammalian tissues have been performed in immortalized cell lines, that are known to be relatively robust. Cells proliferated in vitro in the presence of silicon and “bulk” metabolic assays revealed no toxicity.
There have recently been a number of papers demonstrating the feasibility of interfacing crystalline silicon and mammalian cells (Mayne et al, 2000; Thomas et al, 1999; Curtis and Wilkinson, 1997). The motivation for such studies includes the fact that silicon is easily manipulated into a variety of structures due to developments in the optoelectronics industry and the production of micro electromechanical (MEMS) devices (Steiner and Lang, 1995; Meyer and Biehl, 1995). Starting with crystalline silicon as a substrate, photolithography and etching techniques allow the facile construction of micron- and submicron-sized structures. Silicon surface chemistries targeting the reactivity of silicon oxide via Si—OH groups and silicon hydride (Si—H) have been well explored allowing a variety of surface modifications (Bhatia et al, 1997; Stewart and Buriak, 2000). Other methods for the modification of silicon included electrodeposition machining, laser ablation, laser drilling, micromachining, lithographic galvanic fabrication (LIGA) and embossing. Furthermore, silicon-based cellular arrays can be easily integrated with other silicon-based components such as sensors, heaters, microfluidics arrays, and the like. Porous silicon has recently received considerable interest in applications as a biomaterial due to its solubility in physiologic environments. The primary dissolution product is silicic acid [Si(OH)4] a naturally occurring form of silicon that can be processed and excreted by the body. The rate of dissolution can be controlled by chemical derivatization by methods well known to those skilled in the art (Canham et al, 2000).
The use of crystalline silicon chips as a scaffold for the growth of vascularized perfused microtissue and micro-organ arrays has been taught by Griffith et al. (U.S. Pat. No. 6,197,575, incorporated herein by reference). The apparatus consists of a micromatrix and a perfusion assembly suitable for seeding and attachment of cells on and throughout the matrix and for morphogenesis of seeded cells into complex, hierarchical tissue or organ structures, wherein the matrix includes channels or vessels through which culture medium, blood, gases or other nutrients or body fluids can be perfused. The functional unit in these micromatricies is the channel containing cells and their exudates (such as extracellular matrix molecules) in the desired morphological structure. The channel refers to a hole with defined dimensions, typically 75-1000 micron across, that goes through a sheet of scaffold material approximately 50-500 micron thick. Each channel is sufficiently large to contain a microscale tissue which is a synthetically formed mass of cells forming a tissue structure or a structure that carries out tissue functions. Griffith suggests that such bioreactors would be ideal to simulate liver. One could seed the micromatricies with endothelial cells, followed by the addition of hepatocytes. Alternatively stem cells may be plated directly onto the scaffold and treated with appropriate growth factors to induce differentiation. Such microtissues can be used in the context of an artificial liver apparatus or in drug toxicity and screening assays.
It would be desirable to develop an artificial liver apparatus, similar to a kidney dialysis apparatus, for hepatic support in individuals waiting for liver transplant. However, the liver is a more complex organ than the kidney which is predominantly responsible for salt balance and filtering of molecules based on size. The liver is responsible for detoxification of xenobiotics and hormones, energy metabolism, production of plasma proteins, and production of bile, rather than the simple filtering, of the blood. Furthermore, the factors that lead to hepatic coma in patients suffering from liver failure have not been identified. Thus sustenance of a patient in liver failure with a device that lacks hepatic cells is unlikely. An artificial liver apparatus would need to contain viable, differentiated hepatic cells in order to function. Furthermore, differentiated hepatic cells must be effectively interfaced with the fluid stream, patient plasma, to allow bidirectional mass transfer of large molecular weight proteins. (Allen et al, 2001)
Development of a simulated liver would also be highly desirable for the testing of drugs, both alone in the process of drug development, and to better understand drug interactions (Hodgson, 2001). Initial drug testing is typically performed on cells in culture to facilitate high throughput screening. However, compounds ingested by a patient must have desirable ADMET (absorption, distribution, metabolism, elimination and toxicity) properties in order to be successful as a drug. Such tests can be performed in animals, however there are a number of drawbacks including expense, variation between species, and growing disfavor of the use of animals in research by the general public. As the liver is the initial site of drug metabolism of orally ingested compounds, a number of methods have been developed to simulate the liver in a laboratory setting. These include the use of isolated liver enzymes (e.g. CYP34A and CYP2D6) to assay for conversion or inactivation of drugs. Such a system does not take into consideration that enzymes in the liver do not exist in isolation. Methods using immortalized cell lines can be unreliable. HepG2 cells, a hepatic cell line, have been in culture for 20 years during which time they have ceased to be a truly accurate model of hepatic function. The use of slices of liver from dogs or rats to study bioconversion of compounds are also unsatisfactory due to variations between individual animals and broader species differences. Moreover, a number of compounds alter liver function (e.g. mibefradil, a calcium channel blocker, decreases liver metabolism, leading to the accumulation of cholesterol-lowering drugs). In a system using liver slices, one must run multiple controls on each liver slice to ensure that the data obtained are a result of the compounds being tested rather than variations in liver slices. Therefore, a highly controllable and consistent system that accurately simulates liver function would be useful in developing a better understanding of drug interactions.
A number of strategies have been developed to maintain hepatocytes in a differentiated state in culture. These strategies typically mimic components of the hepatocyte microenvironment in vivo: cell-cell interactions, cell-matrix interactions and soluble cues. For example, heterotypic cell-cell interactions play a fundamental role in liver function. The formation of this vital organ from the endodermal foregut and mesenchymal vascular structures is thought to be mediated by heterotypic interactions. Heterotypic interactions have also been implicated in adult liver physiology (i.e., localization of enzymes in zones of the liver) and pathophysiology (i.e., cirrhosis, and response to injury). In vitro, heterotypic interactions have also proved useful in stabilizing liver-specific functions in isolated hepatocytes. Hepatic cells maintain higher levels of hepatic function at when in contact with non-parenchymal cells as compared to hepatic cells (Bhatia et al, 1999). Seeding cells and maintaining them in such a manner to maintain the optimal ratio of hepatic to non-parenchymal cells is non-trivial. Thus, the maintenance of a culture of differentiated hepatocytes is non-trivial.
Three dimensional bioreactors have been developed to simulate liver function (e.g. U.S. Pat. No. 5,827,729, incorporated herein by reference). In the system, both parenchymal and non-parenchymal cells are seeded onto a porous substrate, preferably a mesh, to form a coculture that will generate tissue in vitro. The mesh is disposed in a container having openings at both ends for media flows. The media flows contain different amounts of nutrients, waste materials, gases and other substances such that a diffusion gradient is established across the tissue. Although such a complex system allows for the growth and maintenance of hepatic cells, such a system is too cumbersome for use in high throughput screening assays.
Cell-matrix contacts are also well-defined in the adult liver, yet are disrupted in standard culture conditions. In vivo, hepatocytes are “sandwiched” by ECM in the Space of Disse. These cell-matrix contacts can be simulated by the use of a “collagen gel sandwich culture.” (U.S. Pat. No. 6,133,030, incorporated herein by reference, and Dunn et al, 1992) Hepatocytes are grown on a support surface and overlaid with collagen. Hepatocytes may be maintained in a differentiated state long term in culture due to the orientation of the ECM interacting with the beta-1 integrins. However, due to the presence of the collagen cell layer as a fragile barrier to bidirectional mass transfer, this culture technique has not been amenable to scale up for clinical bioreactor applications for liver failure.